Since the invention of the scintillation camera by H. O. Anger (H. O Anger, "Scintillation Camera", Rev. Sci. Instr. 29, pp. 27-33, January, 1958 or U.S. Pat. No. 3,011,057) hospital nuclear medicine departments have relied on it for a wide range of human organ imaging. In order to optimize spatial resolution, the modern scintillation camera has evolved to the point where it employs very thin plates of thallium activated sodium iodide (hereinafter termed NaI) as the active scintillation crystal. Typically, these camera plates measure less than 1 cm in thickness, which thereby results in a very small detection efficiency for the available energies of gamma-rays. Because of the low efficiency, most of the radioactive tracer injected into a subject is wasted because the emerging gamma-rays do not interact in the thin crystal. On the other hand, the amount of radioactive tracer that can be injected is limited to a level which minimizes damage caused by gamma-rays interacting in the subject's body. A compromise must therefore be struck. Because of this, it is often necessary to acquire scan data over an extended period of time to obtain enough statistical accuracy to make good quality images. This situation limits the number of scans that can be performed by a single machine in a given period of time. The slow scan time wastes the time of the hospital staff and burdens the subject with an extended period of immobility.
Another contributing reason for the low detection efficiency in standard nuclear medicine systems is the use of a lead collimator between the subject and detector to define the flight direction of the gamma-rays reaching the camera plate crystal. Because gamma-rays are emitted isotropically from the tracer concentrations in the subject's body, a much greater efficiency could be obtained if a wide range of gamma-ray angles could be accepted at the detector. A system which accomplishes this is the positron emission tomograph (PET). In this equipment, the flight direction of each detected emission event is determined by detecting in a second detector the conjugate gamma-ray which is emitted along the opposite direction. Because the PET detectors are uncollimated, a huge increase in detection solid angle is obtained over the scintillation camera.
In addition to the potential advantage of vastly increased sensitivity, a PET detector offers a much greater selection of tracer chemicals. Virtually every biologically interesting element has a positron emitting isotope, whereas single gamma-emitting isotopes are relatively rare in the table of isotopes. PET tracers can be manufactured which are exact chemical copies of metabolically active components of the body, whereas single photon tracers must utilize chemicals such a technetium or thallium which can only partially mimic biological function.
The problem of low detection efficiency in a camera plate is of even greater concern in PET because of high gamma-ray energy (511 keV) and because detectors are required to simultaneously detect two gamma-rays. When two conventional scintillation cameras are used in coincidence for PET the small efficiency of the camera plate is squared, which results in such a small detection efficiency that the gains made by increasing the solid angle are largely cancelled.
In order to avoid these problems, tomographs designed especially for PET typically utilize arrays of small crystals each coupled to a small photomultiplier tube (PMT). These machines obtain fair spatial resolution and good efficiency, but have not been adopted widely for routine clinical studies. Poor acceptance is believed to be partially due to the restrictive detector design. The small crystals are long and narrow and thus it is necessary to employ a rigid ring geometry with the imaged subject volume near the centre of the ring. This rigid geometry creates the need for a large number of crystals to cover the solid angle at a large distance from the subject. Such machines are expensive and inflexible compared to conventional scintillation cameras. In addition, the long narrow crystal shape causes a severe degradation of resolution in regions of the subject away from the ring centre. This distorts the image of small objects so that an object which appears spherical at the centre appears oblong at the edge of the field of view.
A complete three-dimensional fitting procedure has been disclosed under the title, "A Thicker Anger Camera for Gamma-Ray Astronomy", California Institute of Technology, at a conference in November 1984, published IEEE Trans. Nucl. Sci. NS-32, 1985. The procedure by W. R. Cook, M. Finger and T. A. Prince, utilizes a thick crystal Anger camera of otherwise standard configuration. The off-line fitting procedure reported has been shown to improve horizontal position resolution substantially, but does not appear adaptable to on-line event-by-event analysis, especially in high resolution devices. In addition, the procedure described by Cook et al. utilizes a standard NaI camera plate crystal which is inferior to the subject invention in its ability to determine the depth of interaction because the scintillation photons are propagated in an uncontrolled way by the diffuse reflector material surrounding the crystal.
The use of corner reflectors in scintillation cameras has been disclosed in U.S. Pat. No. 3,944,835, Vosburgh, Mar. 16, 1976, but this patent does not recognize the use of corner reflectors to create a depth-independent source distribution. In addition, the corner reflector disclosed by Vosburgh uses a metal material to cause reflection, as is the case with a conventional mirror, rather than a polished surface, to cause total internal reflection.
A system for detecting radiant energy and yielding an output which is indicative of the position of the X, Y plane of the input surface of the detection element is disclosed in U.S. Pat. No. 3,676,676, Somer, July 11, 1972. The Somer system does not relate to scintillation cameras.